The field of the invention is magnetic resonance imaging (“MRI”) methods and systems. More particularly, the invention relates to producing MRI images of blood flowing into moving subjects such as the beating heart without the use of contrast agents.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the excited nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A signal is emitted by the excited nuclei or “spins”, after the excitation signal B1 is terminated, and this signal may be received and processed to form an image.
In MRI systems, the excited spins induce an oscillating sine wave signal in a receiving coil. The frequency of this signal is near the Larmor frequency, and its initial amplitude, A0, is determined by the magnitude of the transverse magnetic moment Mt. The amplitude, A, of the emitted NMR signal decays in an exponential fashion with time, t. The decay constant 1/T*2 depends on the homogeneity of the magnetic field and on T2, which is referred to as the “spin-spin relaxation” constant, or the “transverse relaxation” constant. The T2 constant is inversely proportional to the exponential rate at which the aligned precession of the spins would dephase after removal of the excitation signal B1 in a perfectly homogeneous field. The practical value of the T2 constant is that tissues have different T2 values and this can be exploited as a means of enhancing the contrast between such tissues.
Another important factor that contributes to the amplitude A of the NMR signal is referred to as the spin-lattice relaxation process that is characterized by the time constant T1. It describes the recovery of the net magnetic moment M to its equilibrium value along the axis of magnetic polarization (z). The T1 time constant is longer than T2, much longer in most substances of medical interest. As with the T2 constant, the difference in T1 between tissues can be exploited to provide image contrast.
When utilizing these “MR” signals to produce images, magnetic field gradients (Gx, Gy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available MRI systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically proven pulse sequences and they also enable the development of new pulse sequences.
The MR signals acquired with an MRI system are signal samples of the subject of the examination in Fourier space, or what is often referred to in the art as “k-space”. Each MR measurement cycle, or pulse sequence, typically samples a portion of k-space along a sampling trajectory characteristic of that pulse sequence. Most pulse sequences sample k-space in a roster scan-like pattern sometimes referred to as a “spin-warp”, a “Fourier”, a “rectilinear” or a “Cartesian” scan. The spin-warp scan technique is discussed in an article entitled “Spin-Warp MR Imaging and Applications to Human Whole-Body Imaging” by W. A. Edelstein et al., Physics in Medicine and Biology, Vol. 25, pp. 751-756 (1980). It employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of MR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (2DFT), for example, spatial information is encoded in one direction by applying a phase encoding gradient (Gy) along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient (Gx) in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse Gy is incremented (ΔGY) in the sequence of measurement cycles, or “views” that are acquired during the scan to produce a set of k-space MR data from which an entire image can be reconstructed.
There are many other k-space sampling patterns used by MRI systems These include “radial”, or “projection reconstruction” scans in which k-space is sampled as a set of radial sampling trajectories extending from the center of k-space as described, for example, in U.S. Pat. No. 6,954,067. The pulse sequences for a radial scan are characterized by the lack of a phase encoding gradient and the presence of a readout gradient that changes direction from one pulse sequence view to the next. There are also many k-space sampling methods that are closely related to the radial scan and that sample along a curved k-space sampling trajectory rather than the straight line radial trajectory. Such pulse sequences are described, for example, in “Fast Three Dimensional Sodium Imaging”, MRM, 37:706-715, 1997 by F. E. Boada, et al. and in “Rapid 3D PC-MRA Using Spiral Projection Imaging”, Proc. Intl. Soc. Magn. Reson. Med. 13 (2005) by K. V. Koladia et al and “Spiral Projection Imaging: a new fast 3D trajectory”, Proc. Intl. Soc. Mag. Reson. Med. 13 (2005) by J. G. Pipe and Koladia.
An image is reconstructed from the acquired k-space data by transforming the k-space data set to an image space data set. There are many different methods for performing this task and the method used is often determined by the technique used to acquire the k-space data. With a Cartesian grid of k-space data that results from a 2D or 3D spin-warp acquisition, for example, the most common reconstruction method used is an inverse Fourier transformation (“2DFT” or “3DFT”) along each of the 2 or 3 axes of the data set. With a radial k-space data set and its variations, the most common reconstruction method includes “regridding” the k-space samples to create a Cartesian grid of k-space samples and then perform a 2DFT or 3DFT on the regridded k-space data set. In the alternative, a radial k-space data set can also be transformed to Radon space by performing a 1 DFT of each radial projection view and then transforming the Radon space data set to image space by performing a filtered backprojection.
Magnetic resonance angiography (MRA) uses the magnetic resonance phenomenon to produce images of the human vasculature. To enhance the diagnostic capability of MRA a contrast agent such as gadolinium can be injected into the patient prior to the MRA scan. As described in U.S. Pat. No. 5,417,213 the trick with this contrast enhanced (CE) MRA method is to acquire the central k-space views at the moment the bolus of contrast agent is flowing through the vasculature of interest. Collection of the central lines of k-space during peak arterial enhancement is key to the success of a CEMRA exam. If the central lines of k-space are acquired prior to the arrival of contrast, severe image artifacts can limit the diagnostic information in the image. Alternatively, arterial images acquired after the passage of the peak arterial contrast are sometimes obscured by the enhancement of veins. In many anatomic regions, such as the carotid or renal arteries, the separation between arterial and venous enhancement can be as short as 6 seconds.
When the timing of peak contrast is difficult to perform, dynamic acquisitions are performed in which a series of image frames are acquired. Dynamic acquisitions typically require the use of acquisition sequences of either low spatial resolution or very short repetition times (TR). Short TR acquisition sequences may severely limit the signal-to-noise ratio (SNR) of the acquired images relative to those exams in which longer TRs are possible. One such method using a 3D “Fourier” acquisition is described by Korosec F., Frayne R, Grist T, Mistretta C., “Time-Resolved Contrast-Enhanced 3D MR Angiography”, Magn. Reson. Med. 1996; 36:345-351 and in U.S. Pat. No. 5,713,358. When such a dynamic study is performed the time resolution of the study is determined by how fast the k-space data can be acquired for each image frame. This time resolution objective is often compromised in order to acquire all the k-space data needed to produce image frames of a prescribed resolution without undersampling artifacts.
Efforts have been made to acquire CEMRA images in shorter scan times using undersampled projection reconstruction scanning methods. As described in U.S. Pat. No. 6,487,435, it was discovered that image artifacts due to k-space undersampling are unsubstantial when radial acquisitions are used. This is particularly true of CEMRA image frames in which a pre-contrast mask image is subtracted from each acquired image frame. Further reduction in scan time can be achieved by further undersampling each image frame and reconstructing the image frame using a HYPR reconstruction method as described in U.S. Pat. No. 7,519,412 issued on Apr. 14, 2009.
There are a number of issues presented when a CEMRA method is used. First, repeat imaging is limited because residual contrast agent remains in the patient's vasculature for an extended time. There are also dose concerns when using contrast agents, particularly with subjects having renal insufficiency. Most CEMRA methods also rely on acquiring a pre-contrast mask image and subtraction of the mask from the CEMRA images to suppress background tissues. This subtraction step is problematic when imaging a moving subject such as the beating heart since the two images must be aligned.
An MRA method which does not employ a contrast agent has been developed that encodes spin motion into the phase of the acquired signal as disclosed in U.S. Pat. No. Re. 32,701. These form a class of MRA techniques known as phase contrast (PC) methods. Currently, most PC techniques acquire two images, with each image having a different sensitivity to the same velocity component. Images may then be produced by forming either the phase difference or complex difference between the pair of velocity-encoded images. This motion encoding method is used to image flowing blood in what is commonly referred to as phase contrast magnetic resonance angiography (PCMRA). Phase contrast techniques such as that described in U.S. Pat. No. 6,954,067 have also been used to image flow and provide quantitative measurement of blood flow. In flow imaging the motion encoding gradients used during the scan are sensitive to velocity components in two or three orthogonal directions. From the resulting velocity component images, total quantitative flow images can be produced.
Phase contrast techniques have several shortcomings in a clinical setting. The encoding of motion into the phase of the acquired data requires the application of velocity encoding gradients whose magnitude determines the sensitivity. This velocity encoding gradient setting is a user prescribed parameter (VENC) that is often problematic. A low VENC setting may lead to aliasing of higher velocity spin measurements rendering the directionality of flow difficult to determine. A high VENC setting on the other hand may lead to inadequate sensitivity to low velocity spin motion such as jets that enable the detection of intracardiac shunts. Furthermore, the angle of the prescribed imaging plane needs to be perpendicular to the flow of interest for quantitative analysis or parallel to the frequency direction for qualitative analysis which adds to the technical difficulty of acquisition and produces unusual imaging angles that frequently lead to wrap artifact. Phase contrast techniques also produce complex flow patterns which may be difficult to interpret and difficult to separate from background tissues.
A number of methods have been developed for “tagging” or “labeling” flowing blood prior to flow into a region of interest that do not employ contrast agents or velocity encoding gradients. One such method is referred to as the arterial spin labeling (ASL) family of techniques. These techniques have been used to detect and provide a quantitative measure of blood flow to the brain and depiction of cerebral vasculature as described, for example, by Sallustio et al “Assessment of Intracranial Collateral Flow By Using Dynamic Arterial Spin Labeling MRA and Transcranial Color-Coded Duplex Ultrasound,” Stroke, 2008; 39:1894-97; and by Warmuth et al “Dynamic Spin Labeling Angiography In Extracranial Carotid Artery Stenosis,” AJNR Am J Neuroradiol, 2005; 26:1035-43. In conventional ASL, arterial blood is tagged by magnetic inversion or saturation proximal to a region of interest (ROI) to be imaged. That is, ASL techniques tag blood some distance away from the slice or volume of interest to be imaged. The tagged blood flows into the ROI and the inflow is detected as a modulation of the spin's longitudinal magnetization. This is seen on the reconstructed image as a dark region or a bright region depending on the particular ASL method used.
To create an image of flow, most ASL methods acquire one image with tagged blood and one with untagged (control) blood and subtract the two images. The successfulness of implementing an ASL technique depends, therefore, on accurately aligning the two subtracted images. Such accurate alignment is possible when imaging the head, but problematic when imaging moving subjects such as the beating heart.